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RESEARCH ARTICLE
www.advmatinterfaces.de
Hydrogel-Gated Silicon Nanotransistors for SARS-CoV-2
Antigen Detection in Physiological Ionic Strength
Alexandra Parichenko, Wonyeong Choi, Seonghwan Shin, Marlena Schlecht,
Rafael Gutierrez, Teuku Fawzul Akbar, Carsten Werner, Jeong-Soo Lee,* Bergoi Ibarlucea,*
and Gianaurelio Cuniberti*
The recent Coronavirus Disease 2019 (COVID-19) outbreak strongly propels
advancements in biosensor technology, leading to the emergence of novel
methods for virus detection. Among them, those using nanostructured
field-effect transistors (FETs) provide an ultrasensitive approach toward
point-of-care diagnostics. However, the application of these biosensors in
analyzing biofluids has been limited by their reduced screening length in high
ionic strength liquids. To address this challenge, a solution is presented
involving the surface modification of FETs with a hydrogel based on
star-shaped polyethylene glycol. This hydrogel is loaded with specific
antibodies against the severe acute respiratory syndrome coronavirus
2 (SARS-CoV-2) spike protein. By incorporating the hydrogel, the effective
Debye length is effectively increased, thereby preserving the sensitivity in
biofluids. The efficacy of this approach is demonstrated by employing silicon
nanonet-based FETs for the detection of viral antigens in both buffer and
saliva, as well as cultured viral particle dispersions. Moreover, positive and
negative patient samples are successfully differentiated, showcasing the
practical application of this method. Finally, a theoretical frame is proposed to
elucidate the underlying mechanism behind the preservation of sensitivity.
1. Introduction
The severe acute respiratory syndrome coronavirus 2 (SARS-
CoV-2) pandemic has placed the society in an unprecedented sit-
uation. Although the World Health Organization has reported a
global decline in the number of cases in 2022, these data should
A. Parichenko, R. Gutierrez, B. Ibarlucea, G. Cuniberti
Institute for Materials Science and Max Bergmann Center for
Biomaterials
Dresden University of Technology
Dresden, Germany
E-mail: bergoi.ibarlucea@tu-dresden.de;
gianaurelio.cuniberti@tu-dresden.de
The ORCID identification number(s) for the author(s) of this article
can be found under https://doi.org/./admi.
© The Authors. Advanced Materials Interfaces published by
Wiley-VCH GmbH. This is an open access article under the terms of the
Creative Commons Attribution License, which permits use, distribution
and reproduction in any medium, provided the original work is properly
cited.
DOI: 10.1002/admi.202300391
be analyzed considering the fact that
countries are changing Coronavirus Dis-
ease 2019 (COVID-19) testing strategies.
This shift may result in lowering of
overall number of tests conducted and
consequently, a decrease in the reported
cases.[1] As a result, the true extent of
COVID-19 transmission may be un-
derestimated. Developing a new global
standard testing strategy faster and
more accessible than the most common
polymerase chain reaction (PCR) test
would allow performing more tests and
harmonizing testing procedures world-
wide. Such a standardized approach
would play a pivotal role in early stages
of future pandemics, particularly in sit-
uations where economic constraints are a
concern.
Currently two primary methods are
widely utilized for COVID-19 detec-
tion: the well-established polymerase
chain reaction (PCR) and rapid antigen
tests. The former usually takes from
two to four hours for completion,[2]
which together with the sample delivery to centralized laborato-
ries results in a time-to-result of 24 h. On the other hand, rapid
antigen tests are based on a colorimetric lateral flow assay de-
livering results in few minutes.[3] They are cost-effective, can
be interpreted by naked eye without a dedicated reader, and are
W. Choi, S. Shin, J.-S. Lee
Department of Electrical Engineering
Pohang University of Science and Technology (POSTECH)
Pohang , South Korea
E-mail: ljs@postech.ac.kr
M. Schlecht
Universitätsklinikum Carl Gustav CarusDresden
Dresden, Germany
E-mail: Marlena.Schlecht@ukdd.de
T. F. Akbar,C. Werner
Max Bergmann Center for Biomaterials
Leibniz Institute of Polymer Research Dresden
Dresden, Germany
Adv. Mater. Interfaces 2023, 2300391 (1 of 8) © The Authors. Advanced Materials Interfaces published by Wiley-VCH GmbH
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readily available for individual purchase in many countries. How-
ever, they do not allow quantitative estimation, making it chal-
lenging to monitor the progression of the disease. Moreover, their
sensitivity is highly variable, between 30% and 100%,[4,5 ] show-
ing the best results for individuals showing symptoms and high
viral load.[6]
New miniaturized electronic devices have been introduced
as alternative to the standard diagnostic techniques,[7–10 ] aim-
ing at tackling the aforementioned limitations and toward rapid
detection even in asymptomatic individuals or in those with
low viral load. Nanotechnology has emerged as a valuable tool
in this context. One notable application of nanotechnology in
SARS-CoV-2 detection is the development of breath tests or in-
gestible sensors.[11] The first ones detect volatile organic com-
pounds (VOCs) associated with COVID-19, while the second
ones are designed to be swallowed, allowing them to reach the
gastrointestinal tract where they can detect the presence of in-
flammatory proteins associated with COVID-19.[11] Furthermore,
2D material-based electrochemical and optical biosensors have
shown promise in SARS-CoV-2 detection.[12] These biosensors
utilize ultrathin, 2D materials such as graphene, which pos-
sess unique electrical and optical properties. By functionaliz-
ing these materials with specific receptors or antibodies that
can selectively bind to viral proteins, these biosensors can de-
tect the presence of the virus with high sensitivity. The electri-
cal or optical signals generated by the interaction between the
viral proteins and the nanomaterials can be measured and an-
alyzed to provide accurate diagnostic information.[12] Addition-
ally, nanomaterial-based field-effect transistors (FETs) offer re-
markable ability to detect biorecognition events without the need
of redox mediators or electroactive labels. Various FETs have
been reported for SARS-CoV-2 antigen detection, including those
utilizing organic semiconductors,[13,14 ] carbon nanotubes,[15] or
graphene.[16] Silicon nanowire-based FETs have demonstrated ex-
ceptional sensitivity, as reported by multiple research teams[17,18]
including our group.[19–21 ] Si nanowire FETs have also shown
promise in the realm of SARS-CoV-2 biosensing.[22] However,
a severe limitation of FET-based sensing technology is the De-
bye length screening distance, which dramatically shortens with
the increase of ionic strength of the samples.[23] In physiolog-
ical fluids, this distance is restricted to less than 1 nm, ten-
fold shorter than the size of typical antibodies used as biorecep-
tors. Consequently, pathogen-antibody binding occurs at a dis-
tance from the FET that exceeds the Debye length and thus goes
detected. Despite the introduction of smaller receptors such as
aptamers[24] or nanobodies,[25] antigens still fall outside the de-
tectable area. Further solutions are diluting the samples[26] or
measuring in air[19] to reduce excess of ions, but these approaches
can adversely affect receptor or signal stability. Promising al-
ternatives for direct measurement in high ionic strength have
been found in two main manners: the active control of the De-
bye length via electronics (by electrostatic coupling of multiple
gates[27] or by high frequency sine wave application[28])orvia
surface chemistry. The latter entails the introduction of dielec-
tric polymers, e.g., polyethylene glycol (PEG), on the surface.
This technique has been successfully demonstrated in a variety
of FET types by multiple working groups, showcasing its univer-
sality: silicon nanowires,[29] graphene,[30,31] commercial metal-
oxide-semiconductor FETs (MOSFETs),[32] etc. Typically, PEG is
incorporated as part of a 2D surface modification scheme, form-
ing a mixed monolayer alongside the bioreceptor. Careful con-
trol of the PEG and receptor size, as well as their ratio, is es-
sential. In addition, the system is still subject to disadvantages
inherent to 2D surface modification, with risk of receptor activ-
ity deterioration due to wrong orientation or changes in receptor
conformation, and increased nonspecific interactions.[33] These
problems may be circumvented by integrating bioreceptors in
3D immobilization models such as PEG-based hydrogels, where
the hydrogel is a bioreceptor-hosting platform, simultaneously
increasing the effective Debye length and therefore enhancing
the signal.
PEG-based hydrogels offer a broad range of applications in
the biomedical field due to their biocompatible nature and
superior loading efficiency[34–37 ] of biosensors, the aforemen-
tioned benefits of PEG-hydrogels could be extended to in-
clude the next:[38] protective effects due to antifouling proper-
ties, storage of large bioreceptor amounts in active conforma-
tion, 3D fluid-like environment with higher interaction prob-
ability with target molecules compared to the limited kinetics
of solid–liquid interfaces, filtering of excess of undesired mate-
rial, chemical tailoring to adapt to the specific needs in terms
of bioreceptor nature and transducer material. Previously we in-
vestigated the starPEG-heparin hydrogel on electrodes for im-
pedimetric sensing, employing it as an antibody hosting plat-
form for human immunoglobulin G (IgG) detection.[39] The
portable biosensor was able to detect small antigen concen-
trations down to femtomolar levels in serum. Given that the
size of human IgG is similar to that of SARS-CoV-2 spike pro-
tein (≈10 nm),[40,41 ] it is plausible to expect similar diffusion
kinetics for the latter within the hydrogel. By implementing
this hydrogel on nanomaterial-based FETs, we can further lever-
age the benefits it offers while preserving the high-sensitivity
of FETs through the effective increase of the Debye length. To
our knowledge, these advantages have been anticipated but not
demonstrated, with only a few limited number of studies mon-
itoring local proton content[42] and its variation due to enzy-
matic reactions,[43,44 ] or for DNA detection using hydrogels made
purely of DNA strands which swell upon hybridization with tar-
get sequences.[45]
Here we report the first demonstration of antibody-antigen
binding transduction in high ionic strength using PEG-based
hydrogels as host environment for bioreceptors in conjunction
with nanomaterial-based FETs as transducers (Figure 1). More
specifically, we employ star-shaped PEG to form the hydrogel
matrix, within which antibodies targeting the SARS-CoV-2 spike
protein are hosted. Using silicon nanonet-based FETs, we con-
duct measurements on spiked buffer, spiked saliva, cultured vi-
ral dispersions, and real samples from nasopharyngeal swabs.
Quantitative measuring of virus concentration by this biosen-
sor may offer a valuable means to gain insights into the infec-
tion and disease progression. This data can facilitate the imple-
mentation of early intervention measures. Moreover, the biosen-
sor holds significant potential in preventing the overwhelming
of medical facilities and intensive care units, also in anticipa-
tion of future pandemics. By tailoring the hydrogel composition
according to the requirements of other receptors and biomark-
ers, this technology can be adapted to meet diverse diagnostic
needs.
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Figure 1. Schematic illustration of SARS-Cov- biosensing concept. HM corresponds to heparin maleimide . Created with BioRender.com.
2. Results and Discussion
2.1. Silicon Nanonet FET Modification
The lithographically well-defined semiconducting structure con-
sisted of a silicon nanonet with lateral distances of 100 nm at
the interconnects (Figure 2a). After covering the FETs with the
hydrogel, 3D profiling was conducted using digital microscopy
(Figure 2b). During the imaging process, it was observed that the
hydrogel polymerization resulted in the homogeneous distribu-
tion of the receptor molecule hosting platform. Relative flatness
with a thickness of ≈35 μm was achieved with the application of
pressure with a glass slide during the polymerization (Figure 2b).
The pH sensitivity and therefore the response to variations of
Figure 2. FET characterization. a) Optical microscopy of unmodified FETs and SEM magnification of the nanonet. b) Hydro-gel-modified FET and its D
profile. c) Transfer characteristics and d) pH sensitivity with and without gel.
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Figure 3. Fluorescence intensity through the yellow line presented in the inset (fluorescence microscopy image of an area partially covered with the
fluorescent hydrogel).
charged species was improved after coating (Figure 2c,d). The n-
type FETs showed an increasing current signal proportional to
the applied gate voltage, and dependent on the pH of the buffer
filling the hydrogel. At acidic pH, the increased proton activity
shifted the threshold voltage toward more negative values, with
a pH sensitivity of 28 mV pH−1compared to the 21 mV pH−1of
the unmodified FETs.
In our past work we demonstrated the loading of antibodies in
the starPEG gels by a colorimetric reaction using antibodies la-
beled with the enzyme horseradish peroxidase.[39] This enzyme
oxidizes the transparent substrate tetramethylbenzidine into a
blue-colored solution that can be observed by naked eye. Here
we further verify the hydrogel’s capability to accommodate anti-
bodies by fluorescence imaging. Fluorescently labeled antibodies
were incorporated into the initial mixture, and the resulting gel
was examined using fluorescence microscopy. A significant fluo-
rescence signal was observed in the area coated with the hydrogel
(Figure 3a). The intensity of the signal was measured along the
yellow line indicated in the inset of Figure 3a, which traversed
both the hydrogel-coated and noncoated regions. A noticeable de-
cline in intensity was observed upon crossing the boundary of the
hydrogel coating. The silver transmission lines of the electrodes
contribute to an increase in the intensity of the signal due to re-
flection, therefore we perform the measurement over the silicon
surface. The intensity difference between the hydrogel and the
surface was much lower in the case of gels without antibodies
(Figure 3b).
2.2. Biosensing Response in Spiked Samples
Subsequently, we conducted an analysis of the biosensing re-
sponse in spiked samples. The transfer characteristics of the
FETs were recorded by varying the gate voltage (Vg) while moni-
toring the drain-source current (Ids), at a constant potential (Vds )
of 0.1 V. The signal was measured after each incubation with
increasing concentrations of spike protein receptor binding do-
main (RBD) ranging from 5 pg mL−1to 50 ng mL−1) in phosphate
buffered saline (PBS) (Figure 4a).
The baseline signal was established by recording the signal af-
ter incubation with zero concentration of antigen solution, i.e.,
pure PBS. Following the incubation of the biosensing platform
with different antigen concentrations, the signal exhibited a sig-
nificant shift in its dependence on the gate voltage (Figure 4a,b).
The FETs coated with starPEG-heparin demonstrated a sensitiv-
ity of 30 mV ±5.7 mV to a tenfold increase in RBD concen-
tration. The observed voltage shift toward more positive values
aligns with the theoretical isoelectric point of 3.9 and negative net
charge (z=−1.483) at pH 7.4 for the amino acid sequence of the
RBD (region 480–499: CNGVEGFNCYFPLQSYGFQP). In com-
parison, the biosensor’s performance in diluted buffer showed
a similar sensitivity to antigen concentration changes (31 mV ±
3.5 mV). These results indicate that the hydrogel layer preserves
the sensitivity of the device in high ionic strength solutions. Non-
specific interactions were also tested (Figure 4b,c), where expo-
sure of the hydrogel to human IgG (5 pg mL−1, 500 pg mL−1,
50 ng mL−1) as a protein of high concentration in body fluids re-
sulted in negligible influence compared to incubation with the
antigen.
Purified saliva spiked with RBD was used to confirm the ca-
pability of the biosensor to detect the presence of COVID-19
pathogens in complex biological fluids. Saliva was chosen as po-
tential target fluid due to its ease of obtainment in a less inva-
sive manner compared to a nasopharyngeal swab, while still con-
taining sufficient viral load for detection, as suggested by recent
studies.[46,47 ] As a result, recordings of the I–Vcurves showed de-
pendence on the antigen concentration (Figure 4d). However, the
achieved sensitivity was lower in comparison to that obtained in
PBS (20 mV ±9 mV to tenfold increase of RBD concentration).
This discrepancy can be attributed to the higher viscosity of saliva,
which could lead to a reduced diffusion of the molecules and con-
sequently a diminished interaction between target and biorecep-
tors. Due to the high magnitude of the error bars depicted in the
Figure 4d, the precise quantitative results for a saliva sample with
an unknown viral concentration would be difficult to determine,
rendering the sensor as a semi-quantitative tool. In the future,
the high sensitivity of the biosensor would enable circumventing
the problem by diluting saliva with PBS.
2.3. Biosensing Response with Real Viral Samples
Prior to measuring samples from patients, we tested cultured
virus samples to confirm that the inactivation method resulted in
a detectable RBD content, in order to perform experiments in a
safe environment. Deactivation by 𝛽-propiolactone (BLP) yielded
inconsistent results. BLP is known to alter membrane fusion
during infection, which is attributed to chemical and functional
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Figure 4. Electrical measurements results: a) Biosensor response to SARS-CoV- RBD in PBS, b) Calibration in PBS and in diluted phosphate buer
( ×−) (error bars represent standard deviation of three sensors). Response to spike protein RBD (S) and IgG as nonspecific target are included.
c) Nonspecific interactions of the biosensor with human IgG in PBS, d) Response to RBD in saliva, with calibration graph as inset (error bars represent
standard deviation of three sensors).
modification of proteins involved in the process,[48] including the
spike protein. Large viral aggregates are also formed, hindering
their diffusion through the small pores of the hydrogel.[49] On
the contrary, a heating treatment at 80 °Cfor1h
[50] resulted in
consistent I–Vcurves that showed dependency on virus concen-
tration (Figure 5a). Similar to the measurements with saliva, in-
cubations with highly concentrated samples, even at a fivefold
dilution, resulted in saturation of the biosensing platform. This
observation may be explained by the high virus concentration of
samples obtained by culturing. Nonetheless, all measurements
were conducted in undiluted PBS. Further validation with a heat-
inactivated sample of COVID-19 positive (CT value 15.8) sam-
ple and a negative patient confirmed the capability of the device
to recognize the presence of the virus in a realistic clinical sce-
nario (Figure 5b). However, additional measurements with sam-
ples containing a broader range of viral loads are essential to per-
form a clinical validation of the biosensor. The I–Vcurve recorded
after incubation with the negative sample overlapped with the
baseline originated from a measurement with only PBS, while
incubation with the COVID-19 positive sample resulted in a sig-
nificant shift of 105 mV. Such shift is comparable to the shift cor-
responding to ≈1ngmL
−1of SARS-CoV-2 RBD dissolved in PBS
(Figure 4b) and is close to the shift observed after incubating the
hydrogel layer with an undiluted sample of heat-inactivated cul-
tured virus (Figure 5a). The transfer characteristics of the device
were also obtained for different dilutions of heat inactivated viral
samples from nasopharyngeal swabs (Figure 5c), revealing con-
centration dependency in the signal for all of the three biosensors
used.
To assess the dependence of the biosensor’s sensitivity with
the deactivation technique, an alternative method involving sam-
ple incubation in triton X-100 detergent was tested. This method
resulted in the destruction of the SARS-CoV-2 lipid membrane,
as confirmed by dynamic light scattering (DLS) analysis, which
showed only agglomerations with an average size of 15 nm,
whereas the approximate size of a SARS-CoV-2 particle is 100 nm
in diameter.[41] The voltage shifts obtained with the FETs were
measured for several sample dilutions and compared to those ob-
tained with the heat inactivation method. Both showed a similar
slope with slightly larger shifts observed for the heat inactivation
method (Figure 5d).
Considering the application of hydrogel-modified FETs for di-
rect detection from human samples in a real point-of-care sce-
nario, this new method holds great promise as it allows for de-
tection without the need for any kind of inactivation, where the
spike protein should be readily exposed to antibodies.
The theoretical aspects that qualitatively describe and explain
the preservation of the sensitivity at high ionic strength can be
found in the form of a mathematical model in the Supporting
Information.
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Figure 5. Measurements with viral samples: a) Biosensor response to heat inactivated cultured viral particles in PBS, and a calibration graph as inset
(error bars represent standard deviation of three sensors). b) Transfer characteristics after incubation of the biosensor with COVID- negative (blue)
and COVID- positive (red) heat inactivated nasopharyngeal swabs from patients. c) Response to heat inactivated viral samples from nasopharyngeal
swabs in PBS, and calibration graph (error bars represent standard deviation of three sensors). d) Comparison of the device sensitivity to the SARS-CoV-
viral particles inactivated with two techniques: triton X- inactivation (red), thermal inactivation (grey).
3. Conclusion
In summary, we investigated FETs gated by starPEG-heparin hy-
drogel loaded with antibodies for direct SARS-CoV-2 detection
in high ionic strength solutions. Consecutive incubation of the
hydrogel layer with PBS containing varying concentrations of
RBD sequences (ranging from 5 to 50 pg mL−1) resulted in the
shift of the I–Vcurve toward more positive values after each
incubation–washing cycle. These results are in agreement with
the negatively charged nature of the amino acid sequence at neu-
tral pH and confirm the dependency of the signal on the tar-
get protein concentration. Notably, the device exhibited the abil-
ity to detect the changes in RBD concentration at femtomolar
levels in undiluted buffer, overcoming the Debye length limita-
tion and demonstrating the ultrasensitivity required for early di-
agnostics. Successful results by spiking saliva samples showed
the potential use of this fluid as a noninvasive source, offering
an alternative to the uncomfortable nasopharyngeal swab. Nev-
ertheless, the decreased performance in such viscous fluid sug-
gests that dilutions may be needed to enable accurate direct mea-
surements in human saliva. Furthermore, it is possible to per-
form measurements in real samples containing cultured virus
and distinguish between nasopharyngeal swabs collected from
both healthy individuals and COVID-19 patients. However, in
order to achieve a quantitative result with saliva and real sam-
ples with high viral load, sample dilution would be beneficial
due to the sensor signal saturation. Considering the low de-
tectable concentrations and the additional benefit of an improved
diffusion of molecules using diluted saliva, this would not be
a practical limitation. As a further limitation we may mention
the requirement of incubation–washing cycles. For point-of-care
applications, wash-free assays would be preferable. However, a
rapid immersion of the hydrogel-containing part of the device
into a buffer solution could serve as a viable approach to ensure
obtaining results.
The complexity of the hydrogel poses challenges in precisely
quantifying the Debye length. Nonetheless, the proposed theo-
retical discussion put forth supports the hypothesis that the in-
creased effective Debye length can be attributed to the pegylated
hydrogel. We anticipate that the inclusion of heparin in the gel
can be of great importance for future measurements directly in
blood, owing to its anticoagulant activity.[51] Such measurements
could aid in the identification of other infectious diseases only
detectable in blood or other biomarkers. Further improvement
may be required to prevent saturation of the device following in-
cubation with highly concentrated samples.
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4. Experimental Section
Silicon Nanonet-Based Field-Effect Transistor Fabrication:A silicon-on-
insulator (SOI) wafer with a nm top Si layer (p-type, Ωcm, ())
and a nm buried oxide layer was used as a starting material. The ac-
tive region was formed using a KrF photolithography and an inductively
coupled plasma reactive ion etching (ICP-RIE) and a nm SiOfilm was
grown using a thermal oxidation furnace.
After depositing a photoresist mask on a channel region, arsenic (As)
ions with a dose concentration of × cm−were implanted to form
source and drain regions. Then, the rapid thermal annealing (RTA) was
conducted at °C for s, and a nm oxide was grown as a gate insu-
lator using a pyro-oxidation furnace at °C. Next, contact pads, includ-
ing source/drain transmission lines and the gate electrode, were formed
using an I-line stepper and a lift-o process.
Finally, a SU- passivation layer was formed on the entire surface except
for the contact pads, channel, and reference electrode regions.
Hydrogel Preparation and Deposition:StarPEG-heparin hydrogel was
prepared via Michael addition reaction.[] . mg of heparin maleimide
(HM) were dissolved in μL of PBS. . μL of PBS with dissolved
antibodies against SARS-CoV- RBD epitope (-) with concentration
of the antibodies equal to . μgmL
−were added to . μL PBS solution
containing . mg of PEG-thiol. μL of the HM solution were mixed with
μL of the PBS mixture containing PEG-thiol and the antibodies. The final
mixture was drop casted on top of the FETs for in situ gelation. Due to
the interactions between thiol groups in PEG and maleimide groups of
HM, it polymerized into star-PEG-heparin hydrogel. A glass slide was put
on top immediately after drop casting of the gel ingredients in order to
achieve a homogeneous and relatively flat surface. The hydrogel was left to
further polymerize on top of the FETs for min in humidified atmosphere.
Afterwards the glass slide was removed and the gel was hydrated using
PBS.
Optical Microscopy and Thickness Estimation:Optical microscopy of
the hydrogel layer was done using a digital microscope (VHX-). The
layer of star-PEG-heparin was visualized on top of the FET chip surface.
Scanning mode of the microscope was used to create a D profile and
estimate the thickness of the hydrogel after the deposition procedure de-
scribed above.
Fluorescence Microscopy:Before the imaging procedure, the hydrogel
was prepared following the methodology described above, except the anti-
bodies against RBD. Instead, fluorescently labeled antibodies against hu-
man IgG were mixed with PBS solution containing PEG-SH. Microscopy
of the FET-based biochip covered with the star-PEG-heparin hydrogel was
performed using a fluorescence microscope. The fluorescence intensity
profile was calculated by the Fiji-ImageJ software.
Virus Culture and Deactivation:Virus isolates were obtained from na-
sopharyngeal swabs of anonymous patients. All samples were derived
from anonymous diagnostic patient samples. The use of anonymous ma-
terial for evaluation and validation of methods was approved by the Ethics
Committee of the TU Dresden (BO-EK-). The swab sample was
first filtered through a . ×− filter to avoid contaminations and
then it was added to Vero E cells cultured in DMEM GlutaMAX supple-
mented with % fetal bovine serum, % nonessential amino acids and
% penicillin/streptomycin. The virus was harvested upon destruction of
the cell layer. The supernatant was cleared by centrifugation to remove cell
debris. Passage two of any isolate was used for experiments. Plaque assay
was used to count the viral particles. Briefly, a tenfold dilution series of
the virus stock was used to infect Vero E cells, followed by an overlay-
ing with a semiviscous medium containing .% Avicel to prevent virus
spread through the medium. Three days later, the cells were stained us-
ing % crystal violet in % formaldehyde. The Plaques were counted and
calculated to be ≈×PFU mL−.
Virus stocks were inactivated by adding Triton X- or by heating for
hat°C.
Electrical Measurements:All I–Vcurves with FETs were obtained us-
ing a probe station connected to a source measure unit (B, Keith-
ley Instruments). For the biosensing response, the hydrogel containing
the antibodies against RBD of SARS-CoV- spike protein was incubated
for min with μL of target solution. Every incubation step was fol-
lowed by washing in pure PBS in order to remove unbounded antigens
and other possible molecules or cell debris in the case of detection in
spiked saliva and nasopharyngeal patient samples. Measurements were
taken after each incubation–washing cycle using a drop of PBS ( μL) as
current conducting solution on top of the gate electrode surface covered
with the hydrogel. ×− sodium phosphate buer was used for the
measurements in low ionic strength solutions.
Statistical Analysis:For each sample type (spiked samples and viral
samples), I–Vcurves were obtained from three independent sensors. The
data collected from each sensor were used to calculate the standard de-
viation and average value of the FETs I–Vcurves voltage shifts for each
sample type (normalized to change of gate voltage or ΔVg). The calcula-
tions were done using Microsoft Excel.
Supporting Information
Supporting Information is available from the Wiley Online Library or from
the author.
Acknowledgements
This work was funded by the Sächsische AufbauBank project .
Open access funding enabled and organized by Projekt DEAL.
Conflict of Interest
The authors declare no conflict of interest.
Data Availability Statement
The data that support the findings of this study are available from the cor-
responding author upon reasonable request.
Keywords
COVID- diagnostics, Debye screening length, field-eect transistor , hy-
drogel biosensor, SARS-CoV- detection, silicon nanowires, starPEG
Received: May ,
Revised: June ,
Published online:
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